1. Field of the Invention
This invention relates to surgically implanted metallic prostheses and, in particular, to prosthetic joint implants where one or more metal components are implanted in bone.
2. Description of the Prior Art
Technological advances in prosthetic devices have benefited many people whose joints have become disabled or painful as a result of degenerative diseases, injuries or other causes. The most common and helpful present day implants comprise prostheses to replace all or part of the major weight bearing joints of the body, i.e., the hip joint and the knee joint. Disabilities to these joints severely limit a patient's activities, and thus the development of replacement prostheses for these joints has received primary attention.
Hip and knee joint protheses have been in widespread use in the United States since approximately 1971. Beginning in the early 1970's and continuing thereafter, these prostheses have been implanted using polymethylmethacrylate bone cement. Initially, this approach received widespread acceptance because it almost universally resulted in immediate relief of pain from the diseased joint.
Experience over time however has revealed that a fairly large percentage of the joints implanted with bone cement fail at 5 years and even more at 10 years after implantation. Various explanations have been offered by leading orthopedic surgeons and other experts regarding the cause of these failures. Some experts believe that the failures can be attributed to the body's non-acceptance of bone cement. Others believe that bone cement is well accepted by the body, but is not a proper structural component for use as part of a joint implant because of its physical properties.
Specifically, natural bone has a modulus of elasticity of about 1-4.times.10.sup.6 p.s.i. The metals used for implants generally have a modulus of elasticity on the order of 15-35.times.10.sup.6 p.s.i., that is, the metal has a considerably higher stiffness than the bone. Polymethylmethacrylate cement, on the other hand, has a modulus of elasticity on the order of 0.3-0.5.times.10.sup.6 p.s.i., that is, its stiffness is less than either the metal prosthesis or the surrounding bone. Furthermore, of the three materials--cement, metal, and bone--cement has the lowest mechanical strength and fatigue properties. These physical properties of bone cement in comparison to the physical properties of the natural bone and the metal prosthesis have led many experts to believe that the source of the relatively high failure rate observed for hip and knee prostheses implanted using bone cement is mechanical failure of the cement.
Before the advent of the use of bone cement, prostheses were implanted without cement. These devices sought to achieve fixation by fibrous tissue attachment or by wedging the device into bone. In some instances, the devices included surface features having dimensions on the order of a few millimeters and up so as to try to provide interstices and lattices for engagement with either fibrous tissue or bone. These devices met with varying degrees of success. Perhaps their greatest limitation was that they were not as successful in immediately relieving pain as artificial joints implanted with cement. As a result, during the 1970's, these cementless joints were not widely used.
In the late 1970's and continuing into the 1980's, as the failure rate for cemented prostheses became apparent, interest revived in cementless joints. In particular, with regard to the present invention, efforts were made to develop prosthetic devices whose outer surfaces were porous coated so as to provide an improved interface with natural bone.
As disclosed in Hahn, U.S. Pat. No. 3,605,123, one such effort involved the idea of using plasma flame spraying to coat all or part of the outside surface of a prosthesis with a thin, overlying, porous layer of metal As described in the Hahn patent, the preferred thickness for the layer was from about 0.015 inches to about 0.030 inches, and the pore width at the interface with bone was between 30 microns and 200 microns, with 40 microns to 70 microns considered optimum. As acknowledged in the Hahn patent, overlying a porous layer on a base metal poses a problem in providing a strong bond between the base metal and the overlying porous layer while assuring the provision of an extremely thin layer.
As an alternative to the Hahn system, as disclosed in U.S. Pat. Nos. 3,855,638 and 4,206,516, Robert Pilliar proposed a system in which at least two or three layers of small metal particles were sintered to the outside surface of the prosthesis. As described in the Pilliar patents, the porous coating created by these particles was to have a porosity between about 10 and about 40 percent and an interstitial pore size of more than 50 microns and less than about 200 microns, with the preferred pore size being between about 50 and about 100 microns.
Work through the 1970's with porous coatings established that from about 100 microns to about 500 microns is the most effective range of pore sizes into which bone may grow. This work also led to a consensus among surgeons that the following three elements are needed for a successful implantation of a porous-coated prosthesis: (1) a healthy bone; (2) a precise tight fit of the prosthesis in the cavity created in the bone: and (3) minimum motion between the prosthesis and the bone for some time after implantation to allow at least some of the healing process to take place. To minimize this motion, in the first few post-operative weeks, the patient's activity is much more limited than that permitted with cemented joint implants. With regard to the third element, it was observed that if the device moved within the bone, fibrous tissue, rather than bone, developed at the interface between the prosthesis and the bone. This fibrous tissue attachment can sometimes provide adequate fixation for the prosthesis, but in general is considered less desirable than a direct bone-prosthesis attachment.
Porous coating achieved by either plasma flame spraying or the sintering of small particles raises a number of fundamental concerns regarding the product and its function, including:
(1) The sintering processes can degrade the physical properties of the metal making up the prosthesis. To fuse small particles to the surface of the prosthesis and to one another requires raising the temperature of the prosthesis and the particles to close to their melting temperature. This heat exposure can degrade the physical fatigue properties of the underlying metal. The same is true for plasma flame spraying if a subsequent heating cycle is used to improve the bond between the coating and the substrate.
The degradation in fatigue strength is particularly severe when titanium or titanium alloys are used. See Cook et al., "Fatigue Properties of Carbon- and Porous-Coated Ti-6Al-4V Alloy," Journal of Biomedical Materials Research, 18, 497-512, (1984): Yue et al., "The Fatigue Strength of Porous-Coated Ti-6%Al-4%V Implant Alloy," Journal of Biomedical Materials Research, 18, 1043-1058, (1984): and PCT Patent Publication No. WO 85/03426, published Aug. 15, 1985 and entitled "Apparatus for Affixing a Prosthesis to Bone."
Titanium-containing materials are often preferred for use in prostheses because of their high strength, high degree of biological tolerance by the body, and their greater flexibility in comparison to cobalt-chrome alloys. Specifically, titanium alloys have approximately half the stiffness of cobalt-chrome alloys. Unfortunately, to sinter small particles to titanium-containing materials involves heating the materials to temperatures above their beta transition temperature. This heating transforms the titanium away from its preferred metallurgy and also causes the growth of large grains, which further degrades the physical properties of the metal.
(2) The flame spraying and sintering processes are difficult to control. Specifically, problems arise in achieving strong bonds between the porous coating and the underlying base metal. For example, studies on sintered porous coatings have revealed that for spherical particles having diameters in the range of 100-500 microns, the fixation spots to the substrate metal may be only on the order of 20-30% of the sphere diameter.
As a result of these difficulties in achieving strong bonds, problems arise in ensuring that the adherence of the porous layer to the surface is strong enough to function satisfactorily on the prosthesis, while, at the same time, avoiding physical dimensional changes to the prosthesis as of result of having heated the prosthesis for an extended period of time at high temperatures.
(3) The use of porous surfaces results in a many-fold increase in the area of exposed metal. Although the metals used in prostheses are accepted by the body and are generally considered to be biologically inert, some migration of ions from the metal into the body does take place. Some workers in the art feel that increasing the exposed surface area may increase this migration, and for this reason, believe that porous coating may be undesirable.
(4) In addition to the foregoing, sintering small particles to produce a porous coating results in a structure having various mechanical problems which have not been fully appreciated in the past. The particles used in these processes are commonly small spheres having diameters of between approximately 100 microns and approximately 500 microns. A typical porosity after sintering is on the order of 35%, that is, approximately 35% of the space occupied by the porous coating is available for bone ingrowth. Further, if the process is altered to increase the size of the fixation bond spots, the percentage of porosity is reduced.
Bone and metal have very different strengths, i.e., tensile, compression, and fatigue strengths. Typical strength values for bone are on the order of 1-4.times.10.sup.3 p.s.i.; typical strength values, in particular, fatigue strengths for the metals used in prostheses exceed 50.times.10.sup.3 p.s.i. Thus there is at least a 10 to 1 strength advantage in favor of the metal. Accordingly, since the porous structure serves as a transition zone from metal to bone, it should contain more bone than metal to provide a strength match, e.g., on the order of 10, 20, or at the most 25 percent metal. Yet, with a porosity of only 35%, there is in fact less bone than metal in the existing porous coatings, i.e., just the opposite of what would be desirable from a mechanical point of view.